Drug Delivery Systems and Materials for Wound Healing
Applications
Saghi Saghazadeh1,2,*, Chiara Rinoldi1,2,3,*, Maik Schot1,2,4, Sara Saheb Kashaf1,2,5,
Fatemeh Sharifi1,2,6, Elmira Jalilian1,2, Kristo Nuutila7, Giorgio Giatsidis7, Pooria Mostafalu1,2, Hossein Derakhshandeh8, Kan Yue1,2, Wojciech Swieszkowski3, Adnan
Memic9, Ali Tamayol1,2,8,†, and Ali Khademhosseini1,2,9,10,†
1Biomaterials Innovation Research Center, Department of Medicine, Brigham and Women’s
Hospital, Harvard Medical School. Boston, MA 02139, USA 2Harvard-MIT Division of Health
Sciences and Technology, Massachusetts Institute of Technology. Cambridge, MA 02139, USA
3Materials Design Division, Faculty of Materials Science and Engineering, Warsaw University of
Technology. Warsaw 02-507, Poland 4MIRA Institute of Biomedical Technology and Technical
Medicine, Department of Developmental BioEngineering, University of Twente, Enschede, The
Netherlands 5The University of Chicago Medical Scientist Training Program, Pritzker School of
Medicine, University of Chicago, Chicago, Illinois, USA 6School of Mechanical Engineering, Sharif
University of Technology, Tehran, Iran 7Division of Plastic Surgery, Brigham and Women’s
Hospital, Boston, MA 02115, USA 8Department of Mechanical and Materials Engineering,
University of Nebraska, Lincoln, NE, 68508, USA 9Center of Nanotechnology, Department of
Physics, King Abdulaziz University, Jeddah 21569, Saudi Arabia 10Department of Chemical and
Biomolecular Engineering, Department of Bioengineering, Department of Radiology, California NanoSystems Institute (CNSI), University of California, Los Angeles, CA, 90095, USA
Abstract
Chronic, non-healing wounds place a significant burden on patients and healthcare systems, resulting in impaired mobility, limb amputation, or even death. Chronic wounds result from a disruption in the highly orchestrated cascade of events involved in wound closure. Significant advances in our understanding of the pathophysiology of chronic wounds have resulted in the development of drugs designed to target different aspects of the impaired processes. However, the hostility of the wound environment rich in degradative enzymes and its elevated pH, combined with differences in the time scales of different physiological processes involved in tissue
regeneration require the use of effective drug delivery systems. In this review, we will first discuss the pathophysiology of chronic wounds and then the materials used for engineering drug delivery systems. Different passive and active drug delivery systems used in wound care will be reviewed. In addition, the architecture of the delivery platform and its ability to modulate drug delivery are
†Corresponding authors: A. Tamayol (900 N16th Street, Room NH332; Tel: (402) 472-1601, atamayol@unl.edu); A. Khademhosseini (4121-D Engineering V, Los Angeles, CA 90095-1600, Tel: (617) 388-9271, khademh@ucla.edu).
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Author manuscript
Adv Drug Deliv Rev
. Author manuscript; available in PMC 2019 April 05.Published in final edited form as:
Adv Drug Deliv Rev. 2018 March 01; 127: 138–166. doi:10.1016/j.addr.2018.04.008.
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discussed. Emerging technologies and the opportunities for engineering more effective wound care devices are also highlighted.
Graphical Abstract
1. Introduction
Skin is a barrier protecting internal organs from potential environmental hazards[1]. Skin possesses excellent regenerative properties and injuries or cuts can be healed through a highly orchestrated cascade of physiological events. However, in some cases, this regenerative property is impaired and wounds do not heal in a timely fashion, placing the patients at a significant health risk. Usually, wounds that do not heal in 90 days are referred to as chronic wounds. The treatment of chronic wounds and large burns is expensive and laborious as they are susceptible to infection and often require surgical treatment.
Compromised wound healing exerts a massive burden on the healthcare system. In the US around 4.5 million people need treatment for chronic wounds and it is estimated that over $25 billion is spent annually on management of chronic wounds [2]. Furthermore the burden of chronic wounds is growing due to the increasing incidence of obesity and diabetes [3, 4]. In addition, around 40,000 burn victims are hospitalized every year and 4000 of them die from their injuries. Managing burn wounds is also very challenging due to the extent of the injury [5].
Wound care dates back to several millennia. The most ancient therapies were based on covering the wound with leaves and cloth and applying natural ointments in order to reduce pain, prevent infection, and keep the wound closed. Although some of these strategies are still in practice, they have shown to be insufficient for inducing healing in chronic wounds. In addition, the use of conventional wound care practices for treatment of deep cuts results in the formation of permanent scars. Thus, significant efforts have been dedicated towards developing alternate therapies that restore the regenerative properties of the native skin [6]. These activities can be broadly divided into the following groups: 1) identification of biological processes involved in wound healing and those being disrupted in chronic wounds to find therapeutics that support natural healing mechanisms; 2) development of drug delivery systems that facilitate the effective delivery of therapeutics at the right dosage and time into the wound bed; 3) synthesis of materials that can be used as a scaffold for tissue growth; 4) engineering advanced dressings that function beyond a physical barrier and can sense the wound environments and provide information of its status.
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Drug delivery systems are of particular importance as the ineffective vasculature in wound bed can prevent effective delivery of drug to the healing tissue when the drug is administered systemically. In addition, the side effects of some drugs, the low half-life of biological factors, and the dynamicity of the wound environment require complex drug delivery systems that can deliver the active factors in proper dosage to the appropriate location [7]. Over the past decade, significant progress has been made in the field and many different systems and platforms have been developed.
In this manuscript, the recent progress in various areas of wound care with particular emphasis on drug delivery aspects is critically reviewed. The physiology of wound healing and the pathophysiology of chronic wounds will also be discussed. The materials used for engineering wound dressings and scaffolds for wound care are reviewed and systems designed for controlled release of drugs and factors are discussed. Micro and nano-engineered transdermal drug delivery platforms are also highlighted. A new class of dressings that are smart and can sense the wound environment and can provide information essential for active wound care will also be discussed. The opportunities in the area of drug delivery for effective treatment of wounds will be highlighted. It should be noted that the focus of this review will be on the delivery systems rather than on the delivered therapeutics. 1.1. Wound physiology
Wounds can be categorized into acute and chronic types [8]. Acute wounds are the outcome of traumatic or surgical events that heal predictably following a regular healing process. Burn wounds are another class of wounds that are caused by heat, chemicals, electricity, sunlight, radiation or friction [9]. Burns can be classified into superficial (I°), partial-thickness (II°) and full-partial-thickness burns (III°). I° burns only involve the epidermis and usually heal in a week without any additional procedures. Superficial II° burns undergo re-epithelialization similar to split-thickness skin graft donor sites from dermal appendages in 2 – 3 weeks with good functional and cosmetic outcomes. Deep II° burns require weeks, even months, to re-epithelialize and are associated with prolonged pain and severe scarring. Skin grafting is needed to accelerate the wound coverage. III° burns can heal only to some extent by epithelialization and contraction, however, they usually need skin grafting [10].
Acute skin wound healing is an intricate physiological process that is governed by different cell types, growth factors, chemokines, and cytokines [1]. Traditionally, wound healing processes have been divided into four overlapping phases of hemostasis, inflammation, proliferation, and remodeling that each wound needs to go through in order to heal normally (Figure 1) [11, 12]. The body’s immediate response to injury is hemostasis that occurs at the site of the injury to stop the bleeding and minimize hemorrhage. Platelets and inflammatory cells are the first cells to arrive at the site of the injury by binding to the exposed collagen in the extracellular matrix (ECM). The platelets then secrete a number of proteins such as fibronectin, von Willebrand factor (vWF), sphingosine-1-phosphate, and thrombospondin to enhance further platelet stimulation. The release of clotting factors stimulates fibrin matrix deposition to form a stable clot which serves as a provisional matrix for cells migrating to the wound bed. Also, the aggregation of platelets induces vasoconstriction that reduces blood flow to the wound bed. Platelets trapped inside the clot also secrete other important
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growth factors including transforming growth factors (TGF-α, TGF-β), platelet-derived growth factor (PDGF), insulin growth factor (IGF), interleukin 1 (IL-1) to further progress the subsequent phases of wound healing. TGF-β recruits additional cells, including
neutrophils and macrophages. PDGF also helps with vascularization and recruits fibroblasts, connective tissue cells which deposit collagen and promote repair of the damaged tissue [13].
Inflammation phase occurs immediately after the injury and lasts for about 3 days. During this phase, the complement system activates and together with coagulation generates various vasoactive mediators and chemotactic factors that attract leukocytes to the injury site within the first 24–48 hr after wounding. At this stage, mast cells (a member of leukocytes) release granules filled with enzymes, histamine and other active amines which are mediators responsible for the characteristic signs of inflammation around the wound site. The release of these mediators causes surrounding vessels to become leaky and allows for the efficient movement of neutrophils from the vasculature to the injury site. Fluid accumulation at the wound site causes swelling which is one of the signs of the inflammation [14].
Neutrophils are the next predominant cells in the inflammatory phase that are activated within 24 hr after injury [10]. The major function of the neutrophils is to remove pathogens, foreign material, damaged matrix components and dead cells by the process of phagocytosis. Using different chemical signals, neutrophils are attracted to the site of injury by the process called chemotaxis and attach to endothelial cells in the nearby vessels surrounding the wound. Then, they stimulate endothelial cells to express specialized cell adhesion molecules (CAMs). CAMs function as molecular hooks to recruit more neutrophils to bind to the endothelial cell surface and squeeze through the cell junctions that have been made leaky by the mast cell mediator 4–6 [15, 16]. In about two days after wounding, monocytes and lymphocytes stimulated by cytokines, growth factors, and chemokines migrate to the wound site and differentiate into macrophages that phagocytose remaining necrotic tissue,
pathogens, and debris and initiate the formation of granulation tissue [17]. Therefore, macrophages have a similar function as neutrophils with better regulation of proteolytic destruction of wound tissue through the release of protease inhibitors [18]. During the inflammatory phase macrophages also produce important growth factors (such as TGF-β, PDGF, tumor necrosis factor α, TNF-α) and cytokines (such as IL-1, IL-6) that are responsible for the proliferation of fibroblasts, smooth-muscle cells and endothelial cells as well as and ECM deposition [19]. At the end of the inflammation phase, neutrophils are phagocytosed by macrophages. The reduction in the number of inflammatory cells and factors in the wound indicate the commencement of proliferation phase [13, 20, 21]. The proliferation phase of wound healing is when the wound is “rebuilt” by fibroblast proliferation and collagen deposition to replace the provisional fibrin matrix. The
proliferation phase starts around 2–3 days after the injury and continues until the wound is closed. In this phase angiogenesis, tissue granulation, re-epithelialization, and wound contraction occur [1]. Endothelial cells form new capillaries in a process called angiogenesis that is induced by several growth factors such as VEGF-A, fibroblast growth factor 2 (FGF-2), PDGF, and TGF-β. Angiogenesis is essential for granulation tissue formation during the proliferation phase. Newly formed capillaries bring oxygen and nutrients to the
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growing tissue and take away waste products. The formation of granulation tissue allows the re-epithelialization to begin. The process is stimulated by inflammatory cytokines (such as IL-1 and TNF-a) that stimulate fibroblasts to produce growth factors (such as epidermal growth factor, EGF; keratinocyte growth factor, KGF; and hepatocyte growth factor, HGF) which in turn attracts keratinocytes to migrate to the wound bed [22]. Basal keratinocytes migrate from the wound edges as well as from the skin appendages to the injured area where they proliferate, differentiate and eventually form a cover over the wound. Fibroblasts which have migrated to the wound from bone marrow activate and begin synthesizing the
extracellular matrix by secreting various ECM proteins (such as collagens, fibronectin, and hyaluronan). Macrophages stimulate fibroblasts by secreting PDGF and EGF [23]. The fibroblast already in the wound bed differentiates into specific fibroblasts called
myofibroblasts that close the injured area by pulling the wound edges together in a process called wound contraction [1].
The last phase of wound healing is called maturation and remodeling phase. It begins a couple of weeks after wounding and can last over 1 year. During this phase, all the processes activated in the inflammation and proliferation phases terminate. Endothelial cells,
macrophages, and myofibroblasts that are no longer needed undergo apoptosis or exit the wound [24]. Small capillaries aggregate into larger blood vessels and the metabolic activity of wound healing decrease. The ECM of the damaged area consists mainly of collagen and other ECM proteins. Initial deposition of type III collagen, known as reticular collagen, gradually replaced by type I collagen which is the dominant fibrillar collagen in skin. Fibroblasts secrete the lysyl enzyme oxidase, to realign collagen into an organized network which increases the tensile strength of the tissue to about 80% of normal tissue. This process is orchestrated by various matrix metalloproteinases (MMPs), which fibroblasts and other cells secrete [25, 26]. Migration of cells on ECM and remodeling and degradation of the ECM by MMPs are key elements of wound repair.
1.2. Pathophysiology of chronic wounds
Normally skin repair after wounding is very efficient but under certain conditions wound healing can become impaired. Abnormal wound healing results from a dysfunctional alteration in the carefully regulated biologic processes that characterize normal healing; the restoration of the tissue is either downregulated when the wounds fail to heal or upregulated that results in scarring [27]. The pathophysiology of chronic wounds is still not completely understood but it is known that instead of moving forward in the healing process they get stuck in the inflammation phase. Impaired vascularization and consequent hypoxia, the inability to progress to the healing phase, prolonged and increased inflammation, and inability of immune-cells to control bacterial infection are all critical challenges inhibiting the physiologic healing of chronic wounds [28]. The severe hypoxia creates large regions of avascular/non-viable tissue which is a hospitable environment for bacterial growth and biofilm formation. Biofilm further intensifies inflammation, inhibiting ECM deposition and tissue repair. This condition places the patients in significant danger and recurrent surgical procedures (wound debridement or even tissue amputation) are needed to avoid life-threatening complications.
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Prolonged and overexpression of various interleukins and other inflammatory cytokines (such as TNF-α) prevents the healing process from advancing to the proliferation phase. Hyper inflammation also affects the expression of MMPs that play an important role in wound repair by degrading and removing damaged ECM molecules from the injured tissue. However, their excess proteolytic activity is associated with chronic wounds because they destroy growth factors, cell surface receptors, and temporary ECM essential for cell migration [29, 30]. In addition, lack of growth factors and presence of too many senescent cells in the injured area may result in the inability of these wounds to heal. Inadequate microvasculature can lead to chronic non-healing wounds that are especially common in diabetic patients [28].
Most chronic wounds do not heal through regeneration but through fibrosis forming excessive amounts of connective tissue. Fibrosis also follows chronic inflammation and elevated amounts of pro-inflammatory mediators (such as TGF-β) have been found in the wounds that heal by fibrosis. Growth factor activity is poorly regulated causing unnecessary fibroblast proliferation, neovascularization and increased collagen and fibronectin synthesis [31]. In addition, excessive and prolonged wound contraction occurs resulting in a formation of fibrotic scar tissue [32]. Pathological scars after an injury can be categorized into keloids and hypertrophic scars. Keloids are an abnormal overgrowth of the scar tissue. They extend beyond the boundaries of the original wound and do not regress spontaneously over time. Hypertrophic scars are more common and do not get as big as keloids by not expanding over the borders of the wound. They may also spontaneously regress over time [33].
1.3. Existing wound care systems
As described before, physiologic cutaneous wound healing involves a complex, precisely regulated set of interrelated biological pathways [20, 34–36]. Any factor leading to a variable disruption of these orchestrated phenomena can be a cause of pathologic healing in broader terms. Disrupting factors can be innate (e.g. scarring), exclusively extrinsic (e.g., infections, mechanical stress, et cetera), intrinsic (e.g., aging, diabetic condition, vascular disease, poor systemic conditions, et cetera) or mixed [20, 34–36]. Despite the extremely variable range of conditions that can impair healing, overall we can identify two main scenarios: 1) physiologic healing in a healthy patient with no comorbidities, and 2) pathologic healing in patients with local/systemic comorbidities [2, 3, 20, 29, 34–39]. In healthy patients the goal of ideal wound care therapies would be to 1) protect the wound from external agents (e.g. bacterial infections, mechanical stress), 2) accelerate closure through maintenance of wound moisture, and 3) minimize/avoid scarring. In addition to these goals, wound therapies in patients with local/systemic disorders leading to chronic non-healing wounds would include: 1) removing necrotic tissue and biofilm, 2) modulating inflammation (including edema) and unlocking the inflammatory phase of healing, and 3) boosting the reparative phase of healing (e.g. epithelial migration, granulation tissue formation through collagen deposition and ECM remodeling, angiogenesis, and tissue blood perfusion, and lymphangiogenesis) [2, 3, 20, 29, 34–40]. The TIME (tissue, infection/ inflammation, moisture balance and the edge of the wound) guidelines, proposed in 2002, list some of these factors and try to include them in an integrated therapeutic strategy [41,
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42]. Other elements of wound care that should be taken into consideration to maximize patients’ wellness and outcomes include pain and treatment-related factors such as a decreased frequency of dressing changes, and the reduction in wound-care associated costs [2, 3, 43–45].
Advanced wound dressing should primarily aim to prevent bacterial infection by adequate sealing of the wound micro-environment from external contaminants [46]. The dressing should be able to contain bacterial (or fungal) proliferation, limit infection morbidity and possibly eradicate the pathogen [47, 48]. One of the wound care approaches that have become popular is vacuum assisted closure (VAC) of wound in which a negative pressure is generated at the interface of the wound [49]. In this approach a foam or sponge is placed on the wound and the other side of the porous material gets connected to a vacuum pump, which applies suction. VAC therapy has multiple benefits including: removal of exudates rich with pro-inflammatory cytokines and proteins, reducing the chance of infection and biofilm formation, and increasing the blood supply to the wound bed [50, 51]. Similar to the majority of other wound care products, they are passive and they effectiveness for treatment different types of chronic wounds is not clear. Application of VAC therapy for vast wounds can also be challenging.
Two substantial challenges limit the effectiveness of current wound care strategies: 1) the inability to properly detect bacterial colonization/infection (which is based on clinical inspection and confirmation by laboratory analysis) 2) the non-specificity of the topical antimicrobial/antibacterial therapies [52]. As a consequence, diagnosis of wound infection and administration of therapy are often delayed and are less effective. Misdiagnosed infected wounds are associated with prolonged healing and care, several localized and systemic comorbidities (e.g., amputations in diabetic foot ulcers), and mortality (through bacteremia and septicemia) [53]. Dressings loaded with non-specific antimicrobial drugs have only minimally addressed these challenges. Topical antibacterial/antibiotic therapy has the advantage to reduce effective doses required to kill pathogens and limit systemic effects of drugs [54]. However, unnecessary delivery of antimicrobial/antibiotic substances can be a cause of impaired healing by itself as it can deregulate the cutaneous microbiome, lead the development of antibiotic resistance, and have systemic toxicity. In order to successfully counteract a possible infection, advanced dressings need to provide tightly-regulated patient/ bacteria-specific release of antimicrobial or antibiotic substances [55]. Current wound care strategies mostly rely on the concept of “one treatment fits all” and have been shown to be outdated, ineffective, and non-individualized [34, 44]. In particular, treatments for chronic wound are expensive, labor-intensive, often nonspecific, and rely on a wide range of therapies (including cleansing, debridement, oxygen therapy, antibiotics, and surgery) that are not always integrated to improve therapeutic effectiveness, optimization of medical resources, and patient compliance [56].
Maintenance of physiologic wound moisture and gas exchange is a key feature of effective wound dressings [57, 58]. Several studies have shown that a moist wound environment increases migration and proliferation of keratinocytes, promoting wound epithelialization and closure [59, 60]. It also influences the migration of endothelial cells, angiogenesis, remodeling of the ECM, and it has been associated with a less intense fibrosis [2]. Wound
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moisture can be influenced by several patient-specific factors: while some wounds might present with low levels of hydration others might show excess exudates. Moisture of a wound could also vary over time. Most strategies currently available for wound care adopt a non-specific approach to increase wound moisture [61].
Chronic wounds are often characterized by the presence of high quantities of necrotic tissue which limits healing and is the ideal pabulum for bacterial colonization and growth [8]. Wound debridement can be performed using several different strategies including surgery, autolytic (endogenous), enzymes substances, or other mechanical methods [62]. Several commercially-available dressings provide some embedded debridement capacity.
Incorporating substances capable of effectively removing necrotic tissue in a wound dressing allows for a less painful, continuous, cost-effective, and more physiologic debridement compared to surgery or other methods relying on expensive and invasive medical devices [63].
There are a variety of wound healing products that are currently regulated by the U.S. Food and Drug Administration for use in the United States. These products, mostly in form of dressings, can be categorized as passive, medicated, or interactive dressings [64]. Passive non-drug-eluting dressings have no direct effect on the wound except for acting as a physical barrier. Medicated dressings have been used to promote the healing process either indirectly by removing necrotic tissue, or directly by enhancing wound healing stages. The active agents contained in medicated dressings can include cleaning or debriding agents for necrotic tissue removal, growth factors, antimicrobial agents, and monoterpenes. The dressings used to deliver these agents to wounds include hydrogels, hydrocolloids, alginates, polyurethane foam films, and silicone gels [64]. SANTYL® Ointment is the sole
collagenase-containing biologic debriding agent that has been FDA-approved for treatment of burns and dermal wounds [65].
Wound dressings containing drug(s) are regulated as combination products and are designated to an FDA center that retains the primary regulatory responsibilities and
oversight over the product based on the product’s PMOA. Wound dressings containing drugs are regulated by the Center for Devices and Radiological Health if their PMOA is that of a medical device and otherwise by the drug authorities at the Center for Drug Evaluation and Research [66]. Wound dressings containing drugs are a pre-amendment, unclassified device type that has generally been subject to a premarket review through the 510(k) pathway and cleared for marketing if they are shown to be “substantially equivalent” to a legally marketed predicate device. Wound dressings that do not meet this criterion are automatically Class III under section 513(f)(1) of the FD&C Act. For example, the FDA has mandated that dressings that serve as a replacement for full-thickness skin grafting, accelerate the normal rate of wound healing, or treat full-thickness (3rd degree) burns are Class III medical devices. An example of such a dressing is the Integra® Omnigraft™ Dermal Regeneration Matrix that was approved through the premarket approval (PMA).
In general, existing wound care products have significantly improved the patient care. However, they usually target only one aspect of the impaired cascade and cannot address the multifactorial nature of impaired wound healing. In addition, they are often based on the
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“one product fitting everyone” concept, which significantly reduces their effectiveness. Maintaining the long term activity of fragile molecules and growth factors has been another key challenge preventing their use in commercial products.
2. Materials used in wound care
The materials used in the treatment of chronic wounds are used for two different purposes: 1) scaffolding materials that can host the endogenous cells and facilitate their growth and wound closure; 2) temporary dressings that cover the wound area and maintain a suitable condition supporting the healing process.
The ideal wound dressings are supposed to cover the wound, preserve the body water content, be oxygen permeable to allow oxygen access to growing tissue, and prevent the growth of environmental pathogens without interfering with the wound healing [67]. The utilized materials should be immunocompatible, non-degradable, and should not support cell ingrowth and cellular adhesion so to avoid complications during their removal. Dressing delivering drugs and biological factors should preserve the activity of the drugs and should be able to release the drugs at the desired rate. Another important function of wound dressings is exudate management. Wound exudates contain a large quantity of inflammatory cytokines and chemokines and are a suitable for bacterial growth. The effective removal of wound exudates without dehydrating the tissues is of great importance. The optimal material should guarantee gas and fluid permeability in order to absorb odors, maintain moist conditions and avoid dehydration and exudates accumulation which can result in the formation of necrotic tissue [68].
Scaffolding materials used for the treatment of chronic wounds, should facilitate the tissue regeneration, restore the tissue function, and promote a rapid healing process preventing chronic wounds [69]. The material should possess a degradation rate that matches the rate of tissue growth. In addition, neither the material nor the byproducts of the degradation process should induce immunogenicity and toxicity [69]. The scaffolding material should adhere properly to the surrounding tissues and its mechanical properties should match those of native skin to avoid the detachment and breakage over the course of healing. It should also maintain its water content or strategies should be devised to prevent material dehydration. They should have a limited swelling capacity and maintain their shape over time. These scaffolding materials can also be used as a depot of growth factors and the drug that are directly being delivered to the healing tissue. In this frame, engineered skin substitutes have been explored in order to create a 3 dimensional (3D) architecture that can mimic the ECM and reproduce the natural cell microenvironment [70]. Biomimetic materials are considered the most promising alternative for the production of these constructs. The optimal material should guarantee gas and fluid permeability in order to absorb odors, maintain moist conditions and avoid dehydration and exudates accumulation [68]. In this section, various materials used for engineering wound dressings and skin scaffolds will be discussed and their characteristics will be listed.
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2.1. Materials in temporary dressings
Materials used for wound dressing applications vary in terms of the origin of materials, physical forms, architecture, and properties depending on the specific circumstances. Simply put, there are several different types of wound dressing products in the form of gauze, thin film, foam, hydrogels, hydrocolloids, membranes each of which is suitable for treatment of a specific wound type.
Medical gauze is the most widely used wound dressing product [71]. Gauze is made from woven or nonwoven fabrics based on natural or synthetic fibers, such as cotton yarns and polyester fibers. Gauzes can absorb exudate from the wounds and can keep the environment moist. Moreover, gauzes can be made as sterilized product and can be used in combination with other additives such as petroleum, saline, antibiotics, and antiseptics, or with other wound dressing products, which further expanded their applicability. However, when used alone, gauzes cannot provide good barrier protection against microorganisms. Also, the removal of gauze might cause a second trauma.
Thin film dressing, typically made from polyurethane, is a transparent and elastic synthetic wound dressing product [72]. The elasticity of such polymeric materials allows for comfortable movements of the affected body part. Importantly, polyurethane thin films are semi-permeable, which permits the exchange of oxygen, vapor, and carbon dioxide, but at the same time serves as a barrier to bacteria. Also, the transparency of thin film dressing allows for inspection of wound bed to assess wound healing. However, for wounds with high exudate, the use of polyurethane thin film might lead to accumulation of body fluids and maceration.
Foams made of synthetic polymers such as polyesters also have the mechanical elasticity that accommodates movement of the affected body part and can absorb more exudate than thin film dressing [73]. Moreover, foams are also able to help maintain the moisture environment around the wound bed and permit gas exchange, which are important features to facilitate wound healing. In addition, the porous structures of foams provide cushioning protection of the wounded tissue as well as good thermal isolation properties. Overall, foams are an economic and effective wound dressing that are suitable for various wounds.
However, their high water uptake reduces their effectiveness as drug delivery systems. Hydrogels are crosslinked three-dimensional network structures that can be swollen with a large amount of water [74]. Hydrogel-based wound dressings can be found in amorphous or sheet forms, or as impregnated gauze [73, 75]. Since hydrogels are swollen with water or glycerin, hydrogel wound dressings can donate moisture to dry or minimally exuding wounds. Also, hydrogels can be easily applied to the wound site and can be easily removed when needed. However, the major concern of hydrogel dressings is their permeability to both gas and oxygen, which limits their use against infection. Thus, hydrogel dressing combined with antibacterial compounds has been developed. Kumar et al studied a composite wound dressings made by chitosan hydrogel loaded with ZnO nanoparticles for wound dressings (Figure 2a–c). The antibacterial activity of chitosan and ZnO particles combined with the release of zinc ions improved keratinocyte motility in the wound area and promoted epithelization and healing (Figure 2d) [76]. Another key concern for the use of
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hydrogels is their rapid dehydration without a proper covering. The dehydration, however, might be reduced by incorporation of hygroscopic materials [77].
Another type of dressing is made from sodium and calcium alginate extracted from seaweed. Alginate is a class of natural polysaccharides with mannuronic and guluronic acid units [78]. Alginate has a unique capability of high absorbency of water. As a result, when applied to wounds, the dressing absorbs the exudate to form a hydrogel, which significantly limits bacteria activity. Therefore, alginate dressings are particularly useful for highly exuding wounds. Also, the calcium ions existed in the dressing not only physically crosslinked alginate to form stable physical gels, but also showed bioactivities in certain biological processes involved in wound healing [79]. In one study, the effectiveness of Kaltostat which is a nanofibrous commercially available dressing was compared with two alginate-based dressings crosslinked by the utilization of polyethylene imine and ethylenediamine [80]. The alginate dressings possessed larger pore sizes of about 100–250 μm and facilitated the air permeation (Figure 2e–g). They also better managed the wound moisture and as a result, faster wound healing was achieved in animal studies comparing to the commercial dressing. Hydrocolloid dressings refer to colloidal materials, or gel-forming agents, typically made from gelatin, pectin, or carboxymethylcellulose [81, 82]. When topically applied to wounds, the colloidal materials absorb the exudate to form a gel state that sticks to the wounds and becomes permeable to water and oxygen. Hydrocolloid dressings under working conditions can provide thermal insulation and a moist environment, and are easy to remove due to the lack of mechanical stiffness. Therefore, hydrocolloid dressings are usually produced with a strong and impermeable film backing, which provides isolation of the wounds from bacteria and reduces the chances of infection.
Overall, any of these materials has some advantages, but none of them can be considered as the perfect dressing. Among them, hydrogel dressings that can manage the wound moisture and carry different types of drugs have attracted noticeable attention. The key problem with these dressings is fabricating dressings that fit large burns and skin defects. In addition, the dressing should be designed in a way that could form a conformal contact with the skin and does not constantly move against the healing tissue. The ease of removal is also important and hydrogels and alginate are usually easier to remove without inducing damage in the neo-tissue.
2.2. Biomaterials as skin regeneration scaffolds
The major goals of traditional wound dressing are to protect the wound beds and provide the favorable environment to promote wound healing. However, these products cannot replace the lost tissue, for example, severely damaged dermis [83]. To help repair chronic ulcers that are difficult to heal, cellular tissue-engineered skin substitute products have been developed. These substrates can be seeded with cells to form an engineered skin or can be implanted acellularly to recruit local cells and facilitate their growth. The presence of scaffolds with certain bioactivity can solve the challenge of continuous removal of temporarily ECM caused by excessive MMPs presence. To this end, the use of bioactive materials such as collagen, hyaluronic acid, chitosan has become the focus of current research in this area. Specifically, collagen and hyaluronic acid are the components of the ECM in the living
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tissues, are fully biocompatible and biodegradable, and have demonstrated promising results in vivo [84].
The development of substrates with nanofibrous architectures by the electrospinning technique results in structures that can mimic the natural collagen fibers in ECM [84]. Since the collagen fibers are known to play critical roles in maintaining the integrity and strengths of the skin tissues, electrospun products can provide a scaffold with biomimicking
nanofibrillar structures to promote wound healing. The high surface area of the electrospun fibers allows exudate accumulation, while the interconnected nanopores permit gas permeation. Also, the electrospinning process is compatible with various natural and synthetic polymers. It is possible to use electrospinning tofabricate composite membranes loaded with bioactive species such as antibiotics or nanoparticles to introduce additional functions [85].
Early substrates used as scaffolds for wound care include films, gels, sponges, or membranes based on natural biopolymers such as collagen, chitosan, gelatin, and hyaluronic acid, as well as on synthetic polymers including polyurethanes and polyesters. Composite scaffolds containing both natural and synthetic polymers might provide combined properties such as bioactivity, stability, and mechanical strength. These scaffolds can also be pre-seeded with autologous and allograft cells to facilitate the healing. In this case, the cells either populate and close the wound or serve as factories for production of biological factors regulating the wound environment [86]. Currently, one of the major concerns associated with cell-laden skin substitutes is the relatively high cost for long-term in vitro culture for maturation. To avoid these steps, the development of injectable scaffolds that can encapsulate cells, form tissue constructs in situ, and promote cell migration and organization has been recently demonstrated [87]. Such injectable hydrogels are usually adhesive to the surrounding tissues, fill the wound cavity, and provide a suitable environment for wound closure and promote healing process and tissue regeneration. In vivo studies have demonstrated that full-thickness skin defects treated with antibacterial injectable hydrogels showed thicker granulation tissue and higher collagen deposition compared to commercial skin substitutes, showing the excellence of the injectable systems as candidates for wound healing [88]. Current research efforts in this area are mainly focused on the tuning the adhesion strength and mechanical properties of these injectable scaffolds as well as incorporating biological factors, which can regulate the environment for expedited tissue healing.
2.2.1. Electrospun nanofibrous matrices—Among different scaffolding materials, fibrous substrates have gained noticeable attention for the fabrication of constructs for wound healing. It has been demonstrated that fibrous scaffolds are able to influence cell alignment, shape and function by mimicking the ECM fibrillar organization [89]. Nanofibrous scaffolds can be produced using various techniques such as self-assembly, phase separation, and electrospinning [90]. However, the electrospinning method is the most promising method to fabricate nanofibrous scaffolds [91]. This simple and efficient
technique is based on an electrical field which charges a polymeric solution that is ejected from a syringe and collected on a metallic ground plate [92]. Electrospun nanofibrous scaffolds show high surface to volume ratio, interconnected pores and fiber dimensions in the range of 10–100 nanometers proving that they can properly mimic the ECM native
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structure [93]. Moreover, electrospun matrices show very interesting characteristics such as oxygen-permeability, the possibility of fluid exchange without accumulation (i.e. exudates), suturability and uniform adherence in situ, making them a great candidate for wound healing purposes [94–96]. Electrospun mats of natural proteins such as gelatin and gelatin
methacryloyl (GelMA) have been fabricated and implanted in wound models [97]. The presence of these nanofibrous materials significantly improves the wound healing rate by reduction of necrosis and enhancement of vascularization (Figure 3a–d) [97].
Synthetic polymers such as polycaprolactone (PCL) [93], poly(L-lactic acid) (PLLA) [98], poly(L-lactic acid-co-glycolic acid) (PLGA) [99] are mainly used for producing fibrous scaffolds for skin substitutes. Natural polymers such as gelatin [89], chitosan [100], and collagen [99] were added during the material preparation. Chandrasekaran et al. proposed a biocompatible electrospun scaffold made of poly(L-lactic acid)-co-poly(ε-caprolactone) (PLACL) combined with gelatin [89]. They demonstrate that surface plasma treatment can improve the hydrophilicity, leading to better cell proliferation and collagen deposition [89]. Similarly, chitosan-based constructs have been also investigated for wound healing and dressing applications, in lieu of their strong intermolecular hydrogen bonds, antibacterial activity, and hemostasis [101]. Chitosan-based electrospun scaffolds were demonstrated to support fibroblast viability, adhesion and proliferation in vitro, as well as promote wound healing in a rat model [101].
Other biomimetic compositions have also been electrospun and successfully used for the culture of skin cells [102]. Recently, the encapsulation of bioactive molecules in a synthetic polymeric nanofibrous structure has also been explored in order to improve the
biocompatibility and overcome the low biological properties of synthetic polymers related to the lack of cell-recognition sites [93, 103]. In one example, co-polymers of PCL-PEG were electrospun and then soaked in rhEGF, 1-hydroxybenzole (HOBt), and
1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC) solutions to graft EGF [104]. The results showed superior wound healing in animals receiving functionalized nanofibrous scaffolds (Figure 3e,f) [104]. Anesthetic solutions for pain relief and antibiotics, such as ampicillin [92], for infection treatment have also been integrated into the scaffold structure [92]. In one study, an angiogenic peptide (Vasoactive intestinal peptide, VIP) was
encapsulated in situ as particles over PCL nanofibers. Electrospun nanofibers were first coated with mussel-inspired dopamine, creating an extremely adhesive layer over the nanofibers [105]. The angiogenic peptide (VIP) was then absorbed as a layer over PCL/DA nanofibers. The VIP coated PCL/DA was then immersed in acetone for in situ formation of VIP loaded microspheres within the PCL nanofiber structure. This encapsulation method provided a gradual VIP release over the course of 5 days and the total released amount was significantly higher compared to other samples with microspheres or those without DA component. In vivo application of PCL/DA-VIP nanofibers on full thickness wounds on mice significantly enhanced the wound healing with 96.5% coverage of the wound area at day 7 post surgery. Immunohistochemistry analyses also showed a significant increase in CD31 expression with PCL/DA-VIP group, showing increased angiogenesis [105]. In general, electrospun scaffolds have been shown to preserve the activity of drugs for relatively long periods of time. In addition, depending on their composition, they can
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gradually release the encapsulated drug over a period of days to months. These properties combined with their ECM-like composition make them a suitable choice for scaffolding materials in skin care. However, one key problem that limits their success is their small pore size distribution (smaller than 10 μm) which significantly reduces the rate of cell infiltration and ingrowth. There have been efforts to engineer electrospun scaffolds with larger pore sizes (larger than 20 μm). However, this is an active research area to improve the controllability of electrospinning process [106].
2.2.2. Hydrogel scaffolds—Hydrogels are promising materials for designing scaffolds that promote wound healing [7, 74]. Their intrinsic porous and hydrophilic structure guarantees gas exchange and fluid balance, controlling water evaporation and absorbing exudate and providing moisture to the wound area. Moreover, transparency is an interesting aspect for monitoring the regeneration. It has been demonstrated that hydrogels can sufficiently mimic the ECM structure and functionality, for example by promoting cell adhesion and proliferation and directing cell migration [107]. The hydrogel composition might also influence cell growth, migration, and maturation [108, 109]. Among synthetic and naturally derived hydrogels, as the main components of ECM glycosaminoglycans (GAGs), such as hyaluronic acid and chondroitin sulfate, and collagen-based hydrogel have been studied the most. Researchers demonstrated that the use of GAGs in the hydrogel composition can lead to enhanced cell infiltration, spreading and proliferation. As reported by Kirker et al., hyaluronic acid concentration is one of the key factors for hydrogel resilience and influences cell differentiation and motility [110]. GAG-based hydrogels showed a significantly superior re-epithelialization and higher collagen deposition and organization compared to commercial products [110]. On the other hand, collagen is well known to have proper mechanical and adhesive characteristics but poor angiogenetic properties and fast degradation rate [111]. Therefore, other naturally derived biomaterials such as fibrin, chitosan, dextran and alginate have been explored. Fibrin has been considered for its angiogenic properties and fibrin-based hydrogels have been successful in promoting vascularization and cell recruitment which favors the wound healing process [112].
However, challenging control over their mechanical properties and degradation rate, the risk of immune response or infectious disease transmission, as well as slow crosslinking process limits the successful utilization of fibrin-based scaffolds in wound care [113]. Among naturally derived biomaterials, chitosan has also gained much attention because of its high hydrophilicity which promotes cell adhesion, migration, growth and differentiation. It is well known for its anticholesterolemic, antimicrobial activity and hemostasis which enhance the healing and regeneration process [107]. It has been demonstrated that chitosan-based hydrogels show bactericidal properties if the chitosan concentration is higher than 188 g/mL [107]. In order to guarantee proper angiogenesis Sun et al proposed a dextran-based
hydrogel modified with amine groups to improve adhesion and integration on the wound site. The rapid degradation of the hydrogel structure permitted easy endothelial cell
penetration into the wound area which favored vascularization. It is reported that the system could guide proper tissue regeneration with adequate epidermal morphology [114]. In addition to GAG-based hydrogels, protein-based scaffolds have also been widely used in the literature for promoting wound care. Collagen is the main constituent of ECM and collagen-based scaffolds have successfully supported wound healing. Currently, there are
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commercially available collagen-based scaffolds that are recommended for the treatment of skin disorders. Gelatin is the denatured form of collagen and has been used as scaffolding materials for skin tissue engineering and wound care [115]. Gelatin-based hydrogels adhere to surrounding tissues and it has been demonstrated to support formation of epidermis (Figure 4a,b) [115].
In general, the hydrogel system should also favor the wound healing process and protect it against pathogens. The intrinsic structure should have the potential to support cell growth [112] and encapsulate bioactive molecules such as drugs [116] and growth factors for efficient tissue regeneration. Injectability and in situ hydrogel formation are considered the greatest advantages of the hydrogel structure. The former allows for site-targeted and minimally invasive scaffold implantation during the surgery, minimizing the patient pain and discomfort while the latter permits the obtainment of scaffolds with precise defect shape without fluting or wrinkling in the wound area. Zeng et al. proposed an injectable gelatin hydrogel which provides a suitable and stable environment for adipose derived stem cell loading and trapping maintaining stemness and viability for proper in situ cell delivery [111]. Lately, thermoresponsive hydrogels have also been gaining traction as in situ
formation systems. This typology of hydrogels is very interesting mainly due to their ability of gelation at body temperature. Miguel et al investigated the role of agarose, a marine algae derived biomaterial, in thermosensitive hydrogel compositions. The agarose-based three-dimensional structure presented a rigid network with proper mechanical properties. It has been proven that the system was polymerized in situ at 37°C due to its thermal properties and supported cell adhesion and proliferation [107].
In general, hydrogels are excellent scaffolds supporting tissue ingrowth and eventually wound closure. They offer high water content and ECM mimicking microarchitecture. Their mechanical properties can be tailored to match native tissues and usually can be
functionalized with various proteins. Their relative large pores in comparison to synthetic hydrophobic polymers result in the quick release of freely encapsulated compounds, which reduced their effectiveness as drug delivery tools. Another major shortcoming of hydrogels is their lack of suturability, making their implantation challenging. However, recent advances in the engineering of adhesive hydrogels have somehow solved this major challenge. For example, GelMA and elastin-based adhesives have been developed that are degradable and offer up to 20 times the adhesiveness of commercial fibrin glue [117, 118]. Due to
abundance of collagen and elastin in the Skin ECM, it is expected that these adhesive hydrogels promote skin regeneration. Overall, improving their drug perseverance and ability to release different drugs with suitable release profiles has remained to be addressed. 2.2.3. Foams and spongy scaffolds—Spongy biomaterial structures, obtained by freezing and subsequently freeze-drying of solutions, are already well explored for wound healing purposes. Their large pore size in the range of 50 μm to millimeters allows this kind of system to significantly support cell infiltration, migration and signaling [120]. Porosity and pore size distribution can be controlled by different material concentrations and/or freeze-drying parameters. Due to their high porosity, well-interconnected pores, excellent properties of absorbing fluids and oxygen permeability, sponges have been successfully applied for the treatment of different types of leg ulcers. This system has the great advantage
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of maintaining physiological moist conditions and promoting granulation tissue formation [121]. Mainly naturally-derived biomaterials such as collagen, gelatin, chitosan, and alginate have been developed to produce sponges for wound healing applications. Collagen-based sponges are most commonly used because of their good mechanical and physicochemical properties which prevent wound contraction and promote fluid absorption, respectively. These sponges promote cell adhesion, function, migration, and proliferation of fibroblasts and keratinocytes cultured on their surfaces [121]. Considering the poor antibacterial properties of collagen, Ramanathan proposed a collagen sponge loaded with anti-infective bioactive molecules. Results reported good keratinocyte and fibroblast collagen deposition and growth factor expression and proper re-epithelialization 14 days post application [121]. However, the biggest disadvantages of conventional collagen, generally derived from porcine or bovine sources, are its fast degradation and the risk of human transgenic disease transmission. For these reasons, fish collagen and gelatin have lately gained interest as a potential material for wound dressings. Chandika et al. successfully studied a fish collagen-based sponge scaffold crosslinked with sodium alginate and chito-oligosaccharides for the formation of biocompatible stiff structures with lower biodegradability [122]. Gelatin provides a suitable degradation profile as well as angiogenic properties whilst avoiding the disadvantages associated with collagen. However, its porosity and water uptake
characteristics appeared to be inferior compared to other naturally derived hydrogels made of hyaluronic acid, chitosan or alginate. Gelatin has also often been combined with chitosan which can potentially increase the antibacterial and hemostatic properties of the construct. However, it is reported that the growth factor injection is not adequately efficient due to its high diffusivity and the activity of the biomolecules is not maintained for a long time [123]. In order to overcome these problems, Jinno et al proposed a sponge scaffold composed of 10% acidic gelatin which guarantees the maintenance of FGF-positive charge. They optimized the FGF release rate and the gel composition demonstrating that 7 μg/cm2 could accelerate the wound healing and vascularization [124].
In general, sponges offer larger pore sizes in comparison to the hydrogel scaffolds, which facilitate cellular ingrowth. At the same time, these large pores can potentially affect their mechanical properties and swelling ratio. Thus, the material composition should be engineered in a way that these properties could be controlled over the course of wound healing. The large pores of spongy materials have another negative effect on the quick dispersion of freely loaded drugs and growth factors. Overall, hydrogels have attracted more attention than sponges to engineer scaffolds that can promote wound healing.
2.2.4. Composite materials—As reported above, hydrogels constructs are commonly used as scaffolds for wound healing. However, some unfavorable properties of their intrinsic structures such as their low mechanical strength and inadequate flexibility and inability to allow for long term drug release have limited their use for wound healing applications. Several studies have shown that they do not guarantee wound site protection and they may fail when high cyclic stress is applied. In order overcome these problems, researchers have studied the possibility of incorporating ceramic, metallic and polymeric nanoparticles into both hydrogel and electrospun constructs [125, 126]. The encapsulation of nanoparticles such as zinc oxide, titanium oxide and silver particles as antibacterial agents in the hydrogel
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and electrospun scaffolds has also been explored lately to improve the poor antibacterial properties of the basic structures, avoiding bacterial colonization, local and internal infections and unorganized collagen deposition [127, 128]. Titanium dioxide particles were incorporated in a chitosan hydrogel developed by Behera et al. in order to improve the mechanical and antibacterial properties of the hydrogel structure. The incorporation of the nanoparticles improved the fibroblast attachment, function, spreading and proliferation favoring the wound healing [129]. The incorporation of biodegradable polymeric
nanoparticles which can load, protect and modulate the release of bioactive molecules such as growth factors, drugs and proteins have also been explored lately. In particular, PLGA-based carrier systems are the most exploited due to the lactate degradation products which they release during the degradation process. These micro and nanocarriers will be discussed in details in the following sections. Another advantage of composite systems is improving the adhesion of the engineering constructs. For example, stem rose-mimicking constructs have been generated by electrospinning of branched ZnO particles and PCL (Figure 4c–g) [119]. The generated constructs had slightly larger pore sizes than the pristine polymeric scaffolds, yet offered sufficient mechanical properties. The constructs were extremely potent against bacteria cultures, while supporting the growth of keratinocytes [119].
Overall, the composite systems based on hydrogels or electrospun scaffolds can combine the beneficial properties of the incorporated components. Thus, by combining suitable polymers and drug carriers or biologically relevant micro/nanofeatures, one can engineer scaffolds that meet both the physical and biological requirements for achieving rapid wound healing. The incorporation of drug carriers can also solve the major drawback of hydrogels, which is their insufficient drug release profile.
2.2.5. Bi-layered scaffolds—Chronic wounds, as well as traumatic injuries, usually affect different both dermal and subdermal layers of skin. The dermis possesses low cell density and is composed of ECM deposited and maintained by fibroblasts which support the vascular, lymphatic and nervous systems. Because of its structure, the dermal regeneration is less efficient and more complicated than epithelial regeneration [130]. For these reasons, research groups have recently focused on the development of bi-layered scaffolds which combine both the epithelial and the dermal layers [131]. It has been reported that a system which provides a dense superficial layer and a porous lower layer might be the most promising alternative for complete full-thickness skin regeneration. The epithelial layer should prevent bacterial infiltration and dehydration of the wound site. The ideal dermal layer is instead supposed to have great fluid absorption properties and should promote fibroblast penetration. In this frame, Boucard et al. proposed a novel acellular bi-layered scaffold composed of chitosan hydrogels obtained through a low energy physical
crosslinking method, avoiding any additional chemical agent. The upper hydrogel layer was optimized to be rigid and dense in order to guarantee protection, gas exchange and adequate mechanical properties. A soft lower hydrogel layer was designed to be flexible and able to adapt and adhere to the wound site. A pig animal model was considered for the scaffold implantation evaluation showing collagen I and IV deposition as well as angiogenesis and migration of inflammatory cells. The scaffold promoted the dermal-epidermal interphase regeneration and the wound healing of the full-thickness skin tissue [131].
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The fabrication of multi-layer constructs with suitable properties supporting the growth of different layers can further improve the rate of tissue generation. Such constructs can be fabricated using microfabrication techniques such as molding, lithography, and 3D printing. In addition, if the layers are made from different composite materials, then suitable factors and drugs can be locally released to further enhance the growth of each layer. The benefits and shortcomings of different scaffolds engineered for treatment of wounds are summarized in Table 1.
3. Passive drug delivery systems
The complex process of wound healing involves hemostasis, angiogenesis, and restoring the skin barrier function. The proper occurrence of these processes requires the presence of growth factors and cytokines. However, in some cases these factors are not sufficiently present or are significantly upregulated and may derail the healing process from its normal cascade or completely halt it. Therapeutic agents such as growth factors, cytokines, antibacterial agents, proteins, small molecules, and bioactive agents can improve the rate of physiological processes leading to wound healing. There are several factors that should be considered in deciding the administration route of therapeutics: 1) the dysfunction wound bed vasculature reduces the bioavailability of compounds administered orally or
intravenously; 2) some of the drugs can have systemic side effects; 3) the wound
environment is rich in various pro-inflammatory cytokines that can deactivate the drugs; 4) physiological processes are time consuming and the administered drugs should be present during that time.
Thus, local delivery of therapeutic agents is compelling compared to systemic delivery since it reduces the undesired side effects such as toxicity or suboptimal delivery. Advances in the field of pharmaceutics and micro/nanotechnology have enabled researchers to fabricate drug delivery system that can control the release of the drug in the wound environment or directly deliver the drug to the healing tissue or cells. Localized controlled release provides
spatiotemporal control over the drug dosage at the wound site, protects the drug from metabolic deactivation, and maintains the drug concentration over a prolonged period of time. An optimal drug delivery system should sequentially and selectively release antibacterial agents, growth factors, cytokines, and other small molecules in a controlled way so that the wound would follow the necessary course of healing [132]. The sequential release of therapeutic agents is of paramount importance for chronic wound healing. Chronic wounds suffer from delayed angiogenesis, resulting in extreme hypoxia, followed by a reduction in the production of reactive oxygen species by immune cells. As a result, more pro-inflammatory cytokines are secreted to recruit more immune cells [6]. Continuous infiltration of immune cells without proper healing results in excessive production of pro-inflammatory cytokines such as MMPs, which will excessively degrade the temporary ECM deposited by cells at the injury site and will prevent tissue regeneration. To disrupt the impaired cycle of ischemia, reperfusion, and inflammation, sequential and selective release of anti-inflammatory agents followed by pro-angiogenic growth factors, epidermal growth factors, and small molecules has been suggested. In this section, various drug carriers designed for wound care and technologies used for their fabrication will be reviewed. We will also highlight their release mechanisms.
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3.1. Controlled-release drug carriers
In drug delivery systems, the encapsulated therapeutics can be released through different mechanisms which can be broadly classified as active and passive delivery. In active delivery, the release is triggered in response to environmental stimuli (pH, temperature, enzymes, chemical reactions, redox reactions, etc.) or external stimuli (magnetic field, electric field, light, ultrasound, etc.). In contrast, passive delivery relies on the diffusion of the drug through the carrier matrix to reach the surrounding medium [133]. Inorganic porous drug carriers, such as mesoporous particles, metal-organic frameworks, ceramic or carbon-based nanotubes have been extensively studied as drug carriers since they provide proper encapsulation for poorly soluble drugs and can also protect the drugs from physiological degradation. Organic carriers such as lipid-based systems, layer-by-layer systems, and hydrogels have also been used as passive transdermal drug delivery tools as they are degradable can pass natural epidermal barrier and can be uptaken by targeted cells [134– 136]. In general, carrier size, shape, porosity, degradability, and electrostatic charge can affect the rate and effectiveness of the drug release [137].
Polymeric drug delivery systems are formed from nondegradable or biodegradable
polymers and have been widely used since these systems can be tailored through the physicochemical properties of the polymers as well as various possible encapsulation methods [138]. In polymeric systems, the release is affected by parameters such as molecular weight (Mw), glass transition temperature (Tg), crystallinity, solubility and polymer degradation rate [139, 140]. Polymer molecular weight has a direct effect on the Tg, viscosity, crystallinity, mechanical properties, and degradation rate. Polymers with lower molecular weight have a faster degradation rate and higher elastic modulus. This results in higher deformation and pore expansion upon deformation, leading to higher release. In contrast, polymers with higher molecular weight have lower elastic modulus and are less deformable upon degradation, limiting the drug release [141]. Polymer Tg defines the temperature at which amorphous regions transition from glassy to rubbery state. At
temperatures lower than the Tg, amorphous regions are glassy and have a limited diffusivity. Above the Tg, the amorphous regions have a higher mobility and a significantly higher diffusivity, leading to higher release rates. Since the permeation occurs through amorphous regions, polymer crystallinity is also a detrimental parameter, especially when working with low molecular weight polymers [142, 143]. The hydrophilic/hydrophobic ratio of the polymer also has an impact on the release. Hydrophobic polymeric particles go through surface erosion while hydrophilic polymeric particles swell and the degradation occurs within the bulk of the polymer [143, 144].
Parameters such as polymer chemical composition, molecular weight, and crystallinity degree can modify the degree of polymer solubility in aqueous system. Non-degradable polymeric particles are used to fabricate matrix-type and reservoir-type drug careers. The release mechanism of these systems is mainly diffusion controlled. In matrix-type systems, parameters such as diffusion distance, polymer degree of swelling, and drug concentration gradient determine the diffusion rate. However, in reservoir-type systems, the thickness and permeability of the polymeric particle defines the rate of drug diffusion and release [145]. Drug release from biodegradable polymeric particles occurs by two typical degradation/
